Acoustic transducers for medical ultrasonic imaging are made of piezoelectric materials. A variety of composite materials are made by combining a piezoelectric ceramic with a passive polymer phase. These composite materials extend the range of material properties offered by conventional piezoelectric ceramics and polymers.
In pulse-echo medical ultrasonic imaging, a 1-3 composite geometry has been identified as the most promising. W. A. Smith, “Composite Piezoelectric Materials For Ultrasonic Imaging Transducers—A Review,” 1986 IEEE, CH2358-0/86/0000/0249, pages 249-255. For example, the 1-3 PZT rod-polymer composite structure 10 shown in FIG. 1 consists of a polymer matrix 11 which holds together thin parallel rods 12 of piezoelectric ceramic oriented perpendicular to opposing faces 13, 14 of the plate. Metal electrodes are applied to the faces 13, 14. When a voltage pulse is applied across this plate (in the direction “t”—same as the poling direction) it excites thickness-mode oscillations in the plate in a band of frequencies near the fundamental thickness resonance of the plate. The resulting acoustic vibrations 15 are projected into the soft tissues of the human body where they scatter off organ boundaries and structures within those organs. Echos returning to the transmitting transducer excite thickness oscillations in the piezoelectric plate, which generate an electronic signal used for making an image. By scanning the direction of the interrogating beam and properly interpreting the returning echos, a picture of the interior of the body is produced having substantial diagnostic value to the physician. Smith at p. 249.
Important parameters for a successful piezo material in this application include: sensitivity; acoustical and electrical impedance matching; low electrical and mechanical losses; shapability; thermal stability and structural strength. For good sensitivity, the piezoelectric must efficiently convert between electrical and mechanical energy, so that the electromechanical coupling is high. The piezoelectric must be acoustically matched to the tissue so that the acoustic waves in the transducer and the tissue couple well during both transmission and reception. Each of the array elements electric impedance must be compatible with the driving and receiving electronics, which is usually 50 ohms. For a given geometry of an array element, the electrical impedance is inversely proportional to the dielectric constant of the piezoelectric material. Thus, the dielectric constant must be relatively large. In summary, a good piezoelectric material for medical ultrasonic imaging should have: high electro-mechanical coupling (kt approaching 1); acoustic impedance close to that of the tissue (Z approaching 1.5 Mrayls); reasonably large dielectric constant (∈S≧100); and low electrical (tan δ≦0.10) and mechanical (Qm≧10) losses. See Smith at p. 249.
The performance of a composite piezoceramic varies with the volume fraction of piezoceramic for a given ceramic and polymer. Generally, a trade-off is made between lowering the acoustic impedance and obtaining a high coupling as the volume fraction decreases. Nevertheless, there is a broad range of proportions over which the composite's coupling coefficient is higher and its acoustic impedance lower than those of a pure piezoceramic component. Smith at p. 253.
Transducer arrays have been made from composites, as shown for example in FIG. 2. A composite linear array 20 has rectangular ceramic rods 21 embedded in a polymer matrix 22, with metal electrodes 23, 24 on opposed major surfaces of the composite 27, a matching layer 25 on one major surface for placement adjacent the body, and array elements 26 defined by an electrode pattern on the second major surface. Alternatively, arrays can be made by cutting the composite to isolate array elements. Composites can be made quite flexible, enabling formation into curved shapes for beam focusing and steering.
A second article by W. A. Smith, “New Opportunities In Ultrasonic Transducers Emerging From Innovations In Piezoelectric Materials,” 1992 SPIE International Ultrasonics Symposium (Jul. 21-22, 1992), summarizes the material parameters for various piezoceramic (Table I) and piezopolymer (Table II) materials. Smith also defines a relationship between the three-axis coordinate system and the polar axis of the ceramic, in order to define the independent material parameters (pages 2-3). These relationships define the electromechanical coupling factors, i.e., k31, k33, . . . which measure the true strength of the piezoelectric interaction once the elastic and dielectric response of the medium are normalized out. Known values for the coupling coefficients, as well as the other important material parameters, are listed for some of the major piezoelectric ceramic materials such as barium titanate, lead zirconate titanate, and modified lead titanate, as well as piezopolymers such as polyvinylidene difluoride and its copolymer with trifluoroethylene.
FIG. 3 illustrates schematically three of the various types of composite piezoelectric materials. The composite types are referred to by the connectivity pattern of the individual phases. For example, a 1-3 connectivity refers to a composite with a piezoelectric phase continuous or self-connected in one dimension and a polymer phase self-connected in three dimensions. FIG. 3A shows the 1-3 PZT rods in a polymer structure 30, as previously discussed in regard to FIGS. 1-2. FIG. 3B illustrates a layered 2-2 structure 40, comprising alternating layers of piezoceramic and polymer, wherein electrodes are placed on the opposing top and bottom surfaces. FIG. 3C illustrates a 3-3 composite structure 50, comprising a blend of piezoceramic and polymer. Each of these structures has advantages in different applications. In general, a device structure having a low Q is desired, which is best achieved by efficiently coupling the transducer acoustically to the medium and electrically to the excitation and imaging electronics.
The polymer in each of the FIG. 3 composite structures helps lower the acoustic impedance for a better match with the medium. However, there is still a problem in achieving a good match of electrical impedance. In this regard, it has been proposed to provide a structure of piezoceramic strips interlaced with metal electrodes as shown in FIG. 4B, which is taken from R. Goldberg and S. Smith, “Performance of Multi-Layer 2-D Transducer Arrays,” 1993 Ultrasonic Symposium, 1051-10117-93-0000-1103, IEEE (1993), pages 1103-1106. For comparison purposes, a single layer ceramic element 60 is shown in FIG. 4A, and a multi-layer ceramic element 70 of the same overall dimensions in FIG. 4B, wherein the arrows (61, 71) indicate the poling direction. The stated objective in Goldberg et al. is to use multi-layer ceramics to increase both the transmit and receive sensitivity of a 2-D array element. In the transmit mode, the goal is to increase the acoustic output power into the body tissue for a given source voltage, which is accomplished by matching the electrical impedances of the source and the transducer for maximum power transfer. In the receive mode, the goal is to increase the received voltage that is amplified and processed by the ultrasound imaging system; the received voltage is increased by having a matched transducer impedance relative to the coaxial cable and imaging circuitry. In the Goldberg et al. multilayer structure 70, the ceramic layers 72 (between interlaced electrodes 73) are connected electrically in parallel, and the total clamped capacitance is the sum of the capacitance of each layer. Therefore, the capacitance CN of an N layer transducer with an electrode area A, layer thickness t/N, and dielectric constant ∈ is:CN=N·∈A/(t/N)=N2·Csinglewhere Csingle is the capacitance of a single layer transducer (such as element 60 in FIG. 4A having a single ceramic layer 62 between electrodes 63, 64). As described in Goldberg et al., the open-circuit receive sensitivity is directly proportional to the layer thickness t/N, and as a result increasing the number of layers will decrease the open-circuit sensitivity. However, the authors state that the multilayer ceramic structure's ability to drive an electrical load compensates for the decreased open-circuit sensitivity.
While the multilayer ceramic structure of Goldberg et al. lowers the electrical impedance of an array element for better power transfer with the imaging system, it does not solve the problem of acoustic matching. Furthermore, while 2-D arrays are desirable in providing elements along the azimuth and elevation planes to provide dynamic control of the ultrasound beam in both directions, the smaller size of the array elements increases the electrical impedance, and thus exacerbates the problem of poor transducer sensitivity. Thus, none of the prior art systems effectively provide both a good match of acoustical impedance to the medium being observed, and a good match of electrical impedance of the imaging system, especially for extremely small transducer elements as required in phased arrays and 2-D matrix arrays.
Two-dimensional arrays consist of tiny transducer elements distributed in a square lattice in two dimensions. One of the major problems in 2-D arrays is that element sizes are very tiny which results in extremely large electrical impedance. Even in current phased array elements, the electrical impedance ranges from a couple of hundred ohms to larger than a kilo ohm depending on the frequency and aperture of the elements. In a 2-D array each one of these elements are subdivided into 64 or larger number of elements in the elevation direction. Thus, the impedance of each of the 2-D array elements is at least 64 times larger and makes it difficult to couple the electrical energy from the typically 50 ohm imaging system to the transducer. The present invention solves this electrical impedance problem along with optimizing the acoustic impedance match to the human body.